Photonic blood typing

ABSTRACT

Photonic devices, systems, and methods for detecting an analyte in a biological solution (e.g., whole blood) are provided. Representative photonic devices are optical ring resonators having nanoscale features and micron-sized diameters. Due to the compact size of these devices, many resonators can be disposed on a single substrate and tested simultaneously as a sample is passed over the devices. Typical analytes include blood cells, antibodies, and pathogens, as well as compounds indicative of the presence of blood cells or pathogens (e.g., serology). In certain embodiments, blood type can be determined through photonic sensing using a combination of direct detection of blood cells and serology. By combining the detection signals of multiple devices, the type of blood can be determined.

CROSS-REFERENCES TO RELATED APPLICATIONS

This application is a continuation of U.S. application Ser. No.16/047,911, filed Jul. 27, 2018, which is a continuation of U.S.application Ser. No. 15/436,585, filed Feb. 17, 2017, now U.S. Pat. No.10,073,102, which is a continuation of U.S. application Ser. No.14/234,134, filed May 23, 2014, now U.S. Pat. No. 9,599,613, which is aNational Stage of PCT/US2012/047745, filed Jul. 20, 2012, which claimsthe benefit of U.S. Provisional Application No. 61/588,785, filed Jan.20, 2012, and U.S. Provisional Application No. 61/509,776, filed Jul.20, 2011, the disclosures of which are each expressly incorporatedherein by reference in their entirety.

STATEMENT OF GOVERNMENT LICENSE RIGHTS

This invention was made with government support under 0930411 and1264174, awarded by the national science foundation, underHDTRA1-10-1-0074, awarded by the defense threat reduction agency, andunder N000140910137, awarded by the Office of Naval Research. TheGovernment has certain rights in the invention.

BACKGROUND

Recently, there has been a surge of interest in leveraging siliconphotonics for biosensing, with the ultimate goal of producing a chipwith thousands of orthogonal sensors capable of functioning in clinicalenvironments and available at minimal cost. Silicon photonic biosensorshave already achieved impressive sensitivities relevant for biomedicalapplications. However, the field has been stymied by the challenge ofbiological specificity, the ability to bind preferentially to an analyteof interest when sensing in complex biological samples (e.g., blood,plasma, serum). Here we show, for the first time on a silicon photonicsplatform, label-free biosensing with clinically relevant sensitivity inundiluted human serum. Utilizing a zwitterionic polymer-based surfacechemistry, we dramatically limit the amount of non-specific proteinadsorption to a microring resonator in serum, while maintaining alabel-free sensitivity of 10 ng ml⁻¹. This result represents asignificant step towards the practical application of silicon photonicsfor medical diagnostics and the biomedical sciences.

The noteworthy potential of silicon photonics emerges from thecombination of excellent optical devices with control electronics toproduce inexpensive integrated photonics systems. In recent years,on-chip modulators, detectors, and hybrid lasers have all beendemonstrated. Microring resonators are exceedingly amenable to scalablefabrication using CMOS-compatible processes, setting them apart frommost other resonant optical microcavity devices for high-throughput,multiplexed biosensing. To utilize the microring resonator forbiological sensing, a binding site is chemically introduced on thesurface of the sensor and a shift in the resonance wavelength of themicroring is observed upon analyte binding. Exploiting this approach,SOI microring resonators have been used for the detection of a diverserange of biological species, including antibodies, proteins, nucleicacids, and bacteria, among others.

In evaluating a biosensor for a given application, there are two mainfigures of merit. First, the sensitivity of the device, and second, thecomplexity of the background solution in which the assay can beperformed successfully. Many recent biosensing results have yieldedexcellent sensitivity, but in relatively simple solutions (lowbiological noise). FIG. 1A shows the distribution of attainedsensitivities, as well as the biological noise in each demonstration.The suitability of these platforms for use in clinical assays dependsnot only on their sensitivity, but also on the selectivity of thesesensors for a particular analyte in complex biological fluids (solutionswith high biological noise).

For comparison, FIG. 1A includes the sensitivities of both colorimetricand chemiluminescent enzyme-linked immunosorbent assay (ELISA), thestandard diagnostic technique used by most hospitals. Despite itswidespread use and perceived effectiveness, the ELISA method is notwithout its limitations, as it requires signal amplification of boundanalyte by primary and labeled secondary antibodies, substantiallyincreasing both the cost and time of the diagnostic. A competingtechnology to ELISA, surface plasmon resonance (SPR), has also achievedlow sensitivities in complex media. However, due to the complexity ofplasmonic systems integration, SPR has not been realized as ahighly-parallelized, portable, low-cost clinical assay. Also, while SPRis an effective label-free biosensing technology, the sensing range ofSPR is limited, due to the exponential decay of the surface plasmon fromthe gold substrate. Thus, SPR is poorly suited to study targets, such asbacteria, where the size of the target places the majority of therefractive index change outside of the range of the evanescent wave.Therefore it would be desirable to develop a related label-freetechnology that could be used to sense targets at greater range from thesurface of the device.

Ideally, a diagnostic test should require minimal processing of thebiological sample prior to detection of the analyte of interest.However, non-specific adsorption of proteins in complex biologicalsamples, a process known as fouling, significantly decreases thesensitivity of label-free devices due to a lack of biologicalspecificity at the sensor surface. Common strategies for passivatingsurfaces to non-specific biological interactions include adsorption of‘blocking’ proteins (e.g., serum albumin) and grafting inert polymericscaffolds (e.g., polyethylene glycol) to the surface that increasesurface hydration through intermolecular hydrogen bonding. Thesepassivation strategies are only partially effective and are inadequateto fully resist protein fouling in complex biological samples.

One biological system of great interest is the typing of blood.Immediate bloody typing is not presently enabled by any simple, portabletechnologies. Blood typing is necessary for personalized treatment ofwounds (e.g., in combat situations) and improved safety in bloodbanking.

SUMMARY

This summary is provided to introduce a selection of concepts in asimplified form that are further described below in the DetailedDescription. This summary is not intended to identify key features ofthe claimed subject matter, nor is it intended to be used as an aid indetermining the scope of the claimed subject matter.

In one aspect, a photonic device for determining blood type is provided.In one embodiment, the device includes:

-   -   a sample waveguide having a sample surface; and    -   a binding coating covering and being in optical communication        with at least a portion of the sample surface, the binding        coating being configured to bind a target moiety indicative of        blood type, wherein the photonic device is configured such that        light passed through the sample waveguide has an evanescent        field that extends a distance beyond the sample waveguide        sufficient to detect the target moiety indicative of blood type.

In another aspect, a photonic system for determining a blood type isprovided. In one embodiment, the system includes:

-   -   (1) a first photonic device for determining the blood type,        comprising:        -   a first sample waveguide having a first sample surface; and        -   a first binding coating covering and being in optical            communication with at least a portion of the first sample            surface, the first binding coating being configured to            directly bind to a blood cell target moiety indicative of            the blood type, wherein the blood cell target moiety is            attached to a blood cell body selected from the group            consisting of a blood cell, a blood cell membrane, a blood            cell fragment, a microvesicle, and a blood cell-associated            antigen, and wherein the first photonic device is configured            such that light passed through the first sample waveguide            has an evanescent field that extends a distance beyond the            first sample waveguide sufficient to detect the bound blood            cell target moiety; and    -   (2) a second photonic device for determining blood type,        comprising:        -   a second sample waveguide having a second sample surface;            and        -   a second binding coating covering and being in optical            communication with at least a portion of the second sample            surface, the second binding coating being configured to bind            to an antibody or other acellular biological species            indicative of the blood type, wherein the second photonic            device is configured such that light passed through the            second sample waveguide has an evanescent field that extends            a distance beyond the second sample waveguide sufficient to            detect the antibody indicative of the blood type.

In another aspect, a method for testing a sample to determine a bloodtype is provided. In one embodiment, the method includes:

-   -   (1) applying the sample to a photonic system, comprising:        -   (a) a first photonic device for determining the blood type,            comprising:            -   a first sample waveguide having a first sample surface;                and            -   a first binding coating covering and in optical                communication with at least a portion of the first                sample surface, the first binding coating being                configured to directly bind to a blood cell target                moiety in the sample indicative of the blood type,                wherein the blood cell target moiety is attached to a                blood cell body selected from the group consisting of a                blood cell, a blood cell membrane, a blood cell                fragment, a microvesicle, a microparticle, and a blood                cell-associated antigen, and wherein the first photonic                device is configured such that light passed through the                first sample waveguide has an evanescent field that                extends a distance beyond the first sample waveguide                sufficient to detect the bound blood cell target moiety;                and        -   (b) a second photonic device for determining blood type,            comprising:            -   a second sample waveguide having a second sample                surface; and            -   a second binding coating covering and being in optical                communication with at least a portion of the second                sample surface, the second binding coating comprising                specifically defined antigens configured to bind to an                antibody in the sample indicative of immunity to blood                antigens, wherein the second photonic device is                configured such that light passed through the second                sample waveguide has an evanescent field that extends a                distance beyond the second sample waveguide sufficient                to detect the antibody indicative of immunity to blood                antigens; wherein the first photonic device and the                second photonic device are configured to simultaneously                test the sample; and    -   (2) testing the sample by passing light through the first sample        waveguide and the second sample waveguide.

DESCRIPTION OF THE DRAWINGS

The foregoing aspects and many of the attendant advantages of thisinvention will become more readily appreciated as the same become betterunderstood by reference to the following detailed description, whentaken in conjunction with the accompanying drawings, wherein:

FIGS. 1A-1D. 1A: Distribution of achieved label-free sensitivities forplasmonic, mechanical, and photonic biosensors, as well as the levels ofbiological noise of the solutions tested. For comparison, thesensitivities of colorimetric and chemiluminescent ELISA assays areshown (dashed lines). 1B: Cross-section of the silicon with modalpattern indicated. Contours in |E| are plotted in 10% increments. 1C:Illustration of the biosensor array (left) and a scanning electronmicrograph of a single microring resonator (right). 1D: Schematicillustration of the operation of photonic devices in accordance with thedisclosed embodiments.

FIG. 2. Relative resonance shift for three microrings with varyingsurface chemistries during exposure to undiluted human plasma. Asillustrated, microrings were subjected to buffer, then human plasma, andreturned to buffer. The net resonance shift at the 30 minute mark andbeyond is proportional to the amount of adsorbed material on the surfaceof the microring.

FIGS. 3A-3C. 3A: Resonance peak shifts as a function of time are shownfor representative antiSA-DpC and IgGi-DpC microrings. As illustrated,sensors are initially exposed to buffer (PBS), then SA-spiked buffer (20μg ml⁻¹), and returned to buffer. The insets illustrate the process ofprotein capture on the surface of the antiSA-DpC microring. 3B: Shiftsin resonance as a function of increasing concentrations of SA in buffer.3C: The relative shift difference and a best-fit Langmuir binding curvefor the antiSA-DpC and IgGi-DpC microrings as a function of SAconcentration.

FIGS. 4A and 4B. 4B: The resonance shift as a function of time duringexposure to undiluted human serum spiked with increasing concentrationsof SA. The microrings are washed briefly with buffer (PBS) betweenspiked serum samples. The large shift in all microrings when exposed toserum is expected due to the change in refractive index with respect tobuffer. The relative shift difference is due to the specific binding ofSA to antiSA-DpC sensors. 4A: A detail of the peaks of FIG. 4B.

FIG. 5: Protein fouling on sensors exposed to (1) a solution offibrinogen in PBS (1 mg/ml). Microrings are returned to (2) PBS bufferto observe dissociation of non-specifically adsorbed fibrinogen.BSA-modified microrings are susceptible to fibrinogen fouling, asevidenced by a ˜365 pm shift in the net resonance wavelength of thedevice. However, DpC-modified microrings are highly resistant tofibrinogen adsorption, resulting in no significant overall shift in theresonance wavelength of the microring.

FIG. 6: Representative sensor output for antiSA and IgGi immobilizationto DpC-modified microring resonators. After polymer activation (notshown), sensor baseline is achieved in immobilization buffer followed by(1) solutions of antiSA and IgGi for approximately 15 minutes, prior toa brief wash with (2) immobilization buffer. Residual, unreacted DpCchains are deactivated in (3) a high pH buffer. Microrings are exposedto (4) immobilization buffer for overall quantification of immobilizedantibody.

FIGS. 7A and 7B: The relative shift difference of the microrings as afunction of SA concentration in human serum for antiSA-DpC Microrings 1(FIG. 7A) and 2 (FIG. 7B) follow Langmuir statistics, as expected. Abest fit Langmuir binding relation is included, yielding bindingassociation constants of 0.027 (ng ml⁻¹)⁻¹ and 0.026 (ng ml⁻¹)⁻¹ forMicroring 1 and Microring 2, respectively.

FIG. 8: Response of antiSA-DpC Microring 2 and IgGi-DpC microrings to asecond antiSA solution. After the undiluted human serum/PBS exposuresteps, the microrings are exposed to (1) a solution of a second antiSAfor signal amplification. Microrings are returned to (2) PBS buffer toobserve bound antiSA. As expected, IgGi-DpC cannot bind the secondantiSA probe, confirming that these devices did not bind SA.

FIG. 9: Illustration of hierarchical platform with an ultra low foulingfirst layer and high-loading second layer.

FIGS. 10A-C: 10A: IgG functionalization levels on films with one-layer(One) and hierarchical (HA) structures prepared via SI-ATRP and SI-PIMP.10B: Fouling levels in the presence of undiluted serum or plasma, beforeand after IgG functionalization. 10C: Antigen detection from PBS.

FIG. 11: The SPR sensorgram for the fouling test in the presence ofundiluted blood plasma or serum on hierarchical pCB films.

FIG. 12: Following EDC/NHS activation, the pCB films were modified withTSH IgG, deactivated with SC buffer, and then used for TSH antigendetection.

FIG. 13: Test data for type A red blood cells detected in solution usingring resonator photonic devices in accordance with the embodimentsdisclosed herein.

FIG. 14: Test data for type B red blood cells detected in solution usingring resonator photonic devices in accordance with the embodimentsdisclosed herein.

FIG. 15: Test data for type A and B antigen detected in type A plasmausing ring resonator photonic devices in accordance with the embodimentsdisclosed herein.

FIG. 16: Test data summary for detecting type A and B antigens in type Aand B plasma using ring resonator photonic devices in accordance withthe embodiments disclosed herein.

DETAILED DESCRIPTION

Photonic devices, systems, and methods for detecting an analyte in abiological solution (e.g., whole blood) are provided. Representativephotonic devices are optical ring resonators having nanoscale featuresand micron-sized diameters. Due to the compact size of these devices,many resonators can be disposed on a single substrate and testedsimultaneously as a sample is passed over the devices. Typical analytesinclude blood cells, antibodies, and pathogens, as well as compoundsindicative of the presence of blood cells, antibodies, or pathogens(e.g., serology). In certain embodiments, blood type or blood immunesensitization status can be determined through photonic sensing using acombination of direct detection of blood cells and serology. Bycombining the detection signals of multiple devices, the type of bloodcan be determined.

In one aspect, a photonic device for determining blood type or immunesensitization to blood type is provided. In one embodiment, the deviceincludes:

-   -   a sample waveguide having a sample surface; and    -   a binding coating covering and being in optical communication        with at least a portion of the sample surface, the binding        coating being configured to bind a target moiety indicative of        blood type, wherein the photonic device is configured such that        light passed through the sample waveguide has an evanescent        field that extends a distance beyond the sample waveguide        sufficient to detect the target moiety indicative of blood type.    -   In one embodiment, the binding coating has a first refractive        index prior to binding the target moiety indicative of blood        type and a second refractive index after binding the target        moiety indicative of blood type, wherein the first refractive        index and the second refractive index are different, and wherein        an evanescent field of electromagnetic radiation of a first        wavelength extends beyond the binding coating and into any        binding moiety bound to the binding coating.

FIG. 1D illustrates a representative photonic device 100 used to capturea target moiety 111 from a biological sample fluid flowing over thedevice 100. The target moiety 111 is captured by a binding moiety 109that is part of a binding coating coupled to a surface of a waveguide105 such that the photonic device 100 is operable to detect the boundtarget moiety 111 by evanescent sensing via electromagnetic radiationpropagating through the waveguide 105. As illustrated in FIG. 1B, theevanescent field can extend relatively far beyond the waveguide 105,therefore the photonic device 100 allows for sensing of relatively largespecies (e.g., blood cells) compared with competing techniques (e.g,SPR).

The photonic device 100 includes the optical waveguide 105 disposed on asubstrate 107. The device 100 is configured such that waveguiding willoccur in the waveguide 105 at desired wavelengths of electromagneticradiation. In certain embodiments, the device 100 is surrounded on threesides by the ambient environment (e.g., air) and supported by thesubstrate 107, as illustrated. However, it will be appreciated thatcladding layers can be applied to one or more of the free surfaces ofthe waveguide 105 not abutting the substrate 107.

Still referring to FIG. 1D, a binding moiety 109 is coupled to thesurface of the waveguide 105. The binding moiety 109 can be directlycoupled to the surface of the waveguide 105, or can be attached to anintermediate layer. Such an intermediate layer may be an anti-foulinglayer, as disclosed in the Exemplary Embodiments below.

The binding moiety 109 can be any moiety known to those of skill in theart that will bind to a desired target moiety 111 contained within thesample flow. Representative binding moieties include synthetic orisolated saccharides (mono-, di-, tri-, tetra-, and oligo-)representative of the blood group antigen system, synthetic or isolatedantigenic glycoconjugate (glycopeptide, glycolipid, glycosamino glycan)present on human cells and tissues, synthetic or isolated protein andpeptide moieties and antigens present on human erythrocytes and othercells and tissues, lipid species specific to erythrocytes and othercells and tissues; whole or fragmented eukaryotic/prokaryotic/viralcomponents; synthetic or modified biomimetic compounds capable ofbinding to carbohydrate, protein, glycoconjugate or lipid species;synthetic or isolated saccharides (mono-, di-, tri-, tetra-, and oligo-)antigens or binding moieties; antibodies, nanobodies, fab, aptamers, andother antigen-specific capture species.

Representative binding moieties are configured to bind to: blood cells,antigens, antibodies, pathogens, nucleic acids, and other biologicallyrelevant species. In certain embodiments, the photonic device is used todetermine blood type from a sample of blood in contact with the device.In such embodiments, at least one binding moiety is used to bind atarget moiety is indicative of blood type. Representative targetmoieties include moieties similar to those listed above for the bindingmoieties (e.g., because both direct and indirect typing can be used).Additionally, non-antibody-based capture elements can be used instead ofantibodies, as well as any binding or target moieties known to those ofskill in the art.

In one embodiment, the target moiety indicative of blood type isattached to a blood cell. This approach is referred to as “directtyping”, because the blood cell itself is bound to the photonic device100 (i.e., sensed by the device) through the binding moiety 109 on thewaveguide 105 coupling to the target moiety 111 that is part of theblood cell itself. Blood cells include white blood cells, red bloodcells, platelets, microparticles, and portions thereof.

In other embodiments, “indirect typing” is used, wherein, the bindingcoating is an antigen and wherein the target moiety indicative of bloodtype is an antibody (e.g., through serology) indicative of immunity toblood or pathogen antigens.

While both direct and indirect typing can be used to determine bloodtype, these techniques can also be used to determine what blood typesthe blood is immunized against. Accordingly, indirect typing can be usedto detect antigens that are indicative of immunity to aspects of a bloodtype.

Typically, only one of direct and indirect typing will be used on asingle photonic device (i.e., a single resonator). This is because eachdevice is configured to only bind to one specific target moiety. Thisspecific binding scheme provides certainty that any binding eventdetected by the device will indicate the presence of the single desiredtarget moiety. However, because photonic devices (e.g., ring resonators)can be fabricated to have such small dimensions, a single substrate(“chip” or “die”) may contain up to thousands of devices, each capableof having a different binding moiety attached to its sensing surface.Therefore, on a single substrate, both direct and indirect typing may beused.

This parallel approach proves particularly powerful when undertaking thetyping of blood, which may require sensing of several target moieties(e.g., blood cells and various antibodies) before the blood type can bedetermined accurately. Using known techniques, these multiple targetswould require multiple test runs. Using the present embodiments, asingle blood sample can be typed by contacting it with a singlesubstrate containing a plurality of photonic devices having bindingcoatings configured such that all of the necessary target moieties willbe tested for in the sample. By combining the output of the plurality ofdevices, the target moieties present in the blood can be determined,which, in turn, allows for determination of blood type.

The analysis of blood using both direct and indirect typing is describedfurther below in Exemplary Embodiment 3.

While the devices disclosed have thus far been described as related toblood typing and blood immunology, it will be appreciated that thedevices can also be used to determine the presence of a pathogen in abiological sample, such as blood. As with analysis of blood using thedevices, the pathogen can be detected directly or indirectly byfunctionalizing the device with a binding moiety configured to bind to atarget moiety indicative of a pathogen. The target moiety indicative ofthe pathogen can be one of a pathogen, a pathogen-associated antibody, apathogen associated nucleic acid, and a pathogen-associated antigen.

Binding Coatings

All binding coatings include the binding moiety, as described above.However, the binding coating may also provide other properties. Ofparticular interest are coatings that improve the compatibility of thedevice with the sample. Because biological samples, such as blood, areof particular interest, certain coatings are antifouling, so as todecrease (or eliminate) non-specific binding.

While any antifouling coating can be used, in one embodiment, thebinding coating is zwitterionic. Other antifouling coatings includehydrophilic polymer substrates (e.g. poly- and oligoethylene glycol, PEGand OEG), mono-, oligo- and polysaccharide-based non-fouling coatings,and protein-based coatings that prevent subsequent blood proteinadsorption.

As disclosed in more detail in the Exemplary Embodiments below,zwitterionic films can be configured to provide antifouling propertiesand can be engineered to be both thin (so as to allow for evanescentdetection beyond the film) and to have appropriate binding moieties on adistal surface so as to capture target moieties.

Representative zwitterionic films useful with the embodiments disclosedherein are disclosed in the following references, each of which isincorporated by reference in its entirety: U.S. Pat. No. 7,879,444; U.S.Application Publication Nos. 20110195104, 20090156460, 20100099160,20100247614, 20100249267, 20090259015, 20110097277, 20080181861,20110282005, and 20110105712; and PCT Publication Nos. WO 2009/067562,WO 2008/083390, WO 2009/067566, WO 2009/067565, WO 2008/019381, WO2011/057225, WO 2007/024393, WO 2011/057224, and WO 2011/057219.

In one embodiment, the binding coating comprises a poly(carboxybetaine).However, it will be appreciated that any zwitterionic material can beused as long as it meets the requirements of the devices disclosedherein.

In one embodiment, the binding coating comprises a plurality of layers,including an antifouling layer and a capture layer. As disclosed inExemplary Embodiment 2 (and as illustrated in FIG. 9), the antifoulingcoating can be engineered to have a plurality of layers, each having aseparate function. In one embodiment, the antifouling layer iszwitterionic. In one embodiment, the capture layer comprises at leastone binding moiety. In one embodiment, the binding moiety is configuredto bind to the target moiety indicative of blood type. In certainembodiments, the binding moiety is an antigen configured to capturecirculating antibody to determine indirect type. In another embodiment,the binding moiety is an immobilized antibody or similar antigen-bindingmoiety to determine direct type.

In one embodiment, the antifouling layer is bound to the capture layer.Multilayer binding coatings are typically covalently or ionically boundtogether to provide the antifouling functionality and the bindingfunctionality. However, it will be appreciated that any means forcombining these functionalities is contemplated. Furthermore, in oneembodiment, the binding coating is covalently attached to the samplewaveguide. Conversely, in one embodiment, the binding coating is notbound to the sample waveguide.

Photonic Devices

In one embodiment, the sample waveguide is a portion of a photonicdevice selected from the group consisting of a resonator and aninterferometer.

In certain embodiments, the resonators are optical resonators. Theoptical resonator structures discussed herein can be silicon basednanostructures, and in at least one embodiment include a traveling-wavering that is coupled to a nearby silicon waveguide. In certainembodiments, the ring resonators have a diameter of 100 microns or less.In other embodiments, the ring resonators have a diameter of 50 micronsor less. In at least some embodiments the optical resonator structuresemployed in this technology are nano-scale. In at least someembodiments, optimized slotted waveguides use sub-100 nm features (e.g.,a slot dividing the waveguide in half of 100 nm or less) to concentrateoptical fields near the surface of the waveguides, to achieve relativelygreater sensitivities than can be provided by surface plasmon resonance(SPR) devices. Such slotted ring resonators and waveguide structures areknown in the art.

It should be understood that while a silicon ring resonator structure isan exemplary optical cavity resonator structure that can be used as aphotonic device in accordance with the embodiments disclosed herein, theconcepts disclosed herein are not intended to be limited to siliconstructures or ring resonator structures. The embodiments disclosedherein can be implemented on any photonic devices that can be used forevanescent wave sensing. Furthermore, such photonic can be implementedon varied substrates (not just silicon, as noted immediately above).Thus, where reference is specifically made to silicon ring resonatorsherein, it should be understood that such a structure is an exemplaryand non-limiting embodiment.

Silicon photonics has the potential to revolutionize label-freereal-time bio-sensing. Of particular interest in the present disclosureis the identification of blood type from a sample of whole blood.Through chemistries that can selectively functionalize both oxidizedsilicon and silicon nitride with moieties capable of selectively bindingwith targets (such as proteins, bacteria, and other bio-molecules), itis possible to achieve both specificity and extraordinary sensitivity ina chip-scale system based on the use of nano-photonic waveguides. Theconcepts disclosed herein are based on using the silicon ring resonator,which includes a traveling-wave ring that is coupled to a nearby siliconwaveguide. The ring resonator structure's response is a function of therefractive index (i.e., dielectric constant) above the resonator,permitting it to sensitively and specifically detect bound species(e.g., blood cells, antigens, etc.) at or near the surface of thedevice. Preliminary studies indicate that such ring resonator structurespossess sensitivities that exceed that of SPRs, with limits of detectionlow enough to detect individual small-molecule binding events.

One significance of the use of silicon nano-photonic based devices isthat such silicon-based biosensors can be mass-produced with standardsilicon fabrication techniques widely employed in the electronicsindustry, providing economies of scale enabling powerful yet inexpensivesensing devices to be achieved. As compared to SPR devices, a ringresonator device offers much greater sensitivity, potentially at lowercost, with the possibility of truly integrated data acquisition andprocessing offered by leveraging integrated chip technologies. Byintegrating optical and electronic complexity (photonic waveguides andtransistors) with these sensors, it will become possible to performthousands of different tests, in real-time, on a single sample, with achip that could cost a fraction of the cost of traditional biosensors inlarge volumes. Such a chip would have enormous impact in areas asdisparate as disease diagnosis, global health, biological and chemicalwarfare, homeland security, home health care and diagnosis, andenvironmental monitoring.

In certain embodiments, a single chip device is provided with photonicstructures, switches, detectors, and calibration structures integrateddirectly onto a single die.

Silicon-on-insulator waveguides provide a remarkable platform formanipulating light on a nano-scale. Because silicon-on-insulator is astandard material for manufacturing nano-scale electronic circuits, itis possible to commercially obtain material of extremely high andconsistent quality, and to leverage billions of dollars of commercialnanofabrication infrastructure to build nano-scale devices. Thiscommercial infrastructure also makes silicon an ideal platform formoving rapidly from individual devices into large-scale integratedsystems. Silicon is optically transparent at telecommunicationswavelengths (near 1.5 pm), making the silicon waveguide a system that isinherently compatible with today's existing fiber optic infrastructure.Lastly, silicon has one of the highest refractive indices of any commondielectric material, allowing silicon waveguides to concentrate light toa remarkable degree, in particular near the surfaces of the waveguides,where the field can interact very strongly with surface bound ligands aswell as their targets; viruses, nucleic acids, proteins, and cells.

For evanescent wave sensing, it is extremely desirable to have verylarge optical fields concentrated near the surface of the opticalwaveguides, where binding events can occur. By concentrating an opticalmode into a very small volume, it is a natural corollary that the peakelectric field strength of the optical mode will increase. It ispossible, with integrated optics in a high index contrast system likesilicon ridge waveguides, to achieve mode field concentrations that are10,000× or more what is typically achieved for a propagating(non-focused) beam of light in air. In fact, the electric field strengthof the optical mode propagating in a typical nano-scale siliconwaveguide is comparable to the concentrations that can be achieved atthe focus of a tightly converging beam in free space. With a siliconguide, this mode can propagate for centimeters without substantiallosses, whereas the focus of a lens is only microns in length. Incomparing the sensitivity of a silicon photonic system to a conventionalSPR biosensor solution, it is worth considering the relevant pathlengths along which light can interact with the ligand molecules. In anSPR system, the light bounces through the thin (˜10 nm) layer of ligandonly twice. By contrast, with ring resonators, the light will travelaround an approximately 50 pm diameter ring many thousands of times onaverage (or more), providing a radically increased interaction length.

An additional advantage comes from the ability to fabricate many opticaldevices within the same chip, and to use lithography to align themtogether. As a result, there is only one alignment needed in packagingthe devices; it is possible to address an entire optical system with asingle fiber array connecting to the outside world. By integratingmultiple devices onto the same chip, a single optical alignment can beused to address hundreds or even thousands of different opticalcomponents, all of which can comprise a single complex system.

Such a system could include on-chip resonators, detectors, a switchmatrix, couplers, and transistor-based control electronics. This meansthat once a system is in place to test and package these integrateddevices, the marginal cost of adding more complexity to a given deviceis very small.

One of the great benefits of working in the silicon-on-insulator systemfor photonics comes from access to cutting-edge lithographic processes.By etching an extremely narrow trench (5-100 nm) down the center of awaveguide, it becomes possible to confine a significant fraction of thepropagating optical mode in the low-index slot formed in the center ofthe guide. The divergence condition for a transverse electric mode,moreover, causes the optical field to be concentrated in the low-indexregion between the slots.

In certain embodiments, the silicon based ring resonator structure andfluidics components are implemented on an integrated silicon chip. Acontroller can be implemented as a custom designed circuit (such as anapplication specific integrated circuit) or a microprocessor and memory,the memory including machine instructions which when executed implementa plurality of functions, including introducing a sample into the ringresonator structure, collecting optical data related to the resonance ofthe ring structure, and analyzing the results to determine if aparticular target moiety has been bound to the ring resonator structure.

In certain embodiments, fluidics components are designed to establishspecific conditions in the sampling volume of the ring resonatorstructure to facilitate the study of the sample volume. The sample maybe introduced to one or more devices using fluidics. In such anembodiment, the fluidics components include a configurable micro-fluidicflow cell in which flow can be precisely controlled. Such fluidiccomponents include flow cells that can be incorporated into a siliconbased chip. Incorporating micro-fluidics with the photonic devicesenable simplified sample delivery, routing, and analysis as related tothe photonic devices.

There are many different resonator structures that can be implemented onsilicon or silicon related substrates, which can achieve evanescent wavecoupling effects at significantly greater distances from a sensorsurface than can SPR devices. Such resonator structures, one of which isthe silicone ring resonator structure discussed above, include photonicresonator crystals, photonic resonator rings, photonic resonator disks,photonic resonator linear cavities, photonic resonator racetracks,photonic distributed Brag reflectors, and Fabry-Pérot structures. Suchstructures are encompassed by the term optical micro cavity resonators(or optical cavity resonators). These high-index-contrast optical cavityresonators can be implemented using silicon, silicon nitride, germanium,or any mixture thereof, in crystalline, polycrystalline and amorphousforms.

In another aspect, a photonic system for determining a blood type isprovided. In one embodiment, the system includes:

-   -   (1) a first photonic device for determining the blood type,        comprising:        -   a first sample waveguide having a first sample surface; and        -   a first binding coating covering and being in optical            communication with at least a portion of the first sample            surface, the first binding coating being configured to            directly bind to a blood cell moiety indicative of the blood            type, wherein the blood cell moiety is attached to a blood            cell body selected from the group consisting of a blood            cell, a blood cell membrane, a blood cell fragment, a            microvesicle, and a blood cell-associated antigen, and            wherein the first photonic device is configured such that            light passed through the first sample waveguide has an            evanescent field that extends a distance beyond the first            sample waveguide sufficient to detect the bound blood cell            moiety; and    -   (2) a second photonic device for determining blood type,        comprising:        -   a second sample waveguide having a second sample surface;            and        -   a second binding coating covering and being in optical            communication with at least a portion of the second sample            surface, the second binding coating being configured to bind            to an antibody indicative of the blood type, wherein the            second photonic device is configured such that light passed            through the second sample waveguide has an evanescent field            that extends a distance beyond the second sample waveguide            sufficient to detect the antibody indicative of the blood            type.

In one embodiment, the first photonic device and the second photonicdevice are configured to simultaneously determine the blood type. Byusing simultaneous detection, the output of multiple devices can becompared to determine blood type (or detect a pathogen) in real time.

In one embodiment, the photonic system further comprises a computerconfigured to determine the blood type using output from both the firstphotonic device and the second photonic device. By comparing theresponse of the two different devices having two different bindingcoatings, a greater depth of information about the sample can beobtained using a single system. For example, both blood type (directbinding) and serology can be performed at the same time on the samesample.

In one embodiment, the photonic system further comprises a referencewaveguide that does not have any binding coating. The referencewaveguide may be an untreated waveguide that serves as a temperaturereference that allows for calibration of the obtained data with regardto the temperature of the devices (e.g., resonators changecharacteristics with changing temperature and so fluctuations intemperature should be accounted for in interpreting data from devices).

In one embodiment, the photonic system further comprises a thirdphotonic device configured to bind a moiety indicative of a pathogen.Using yet another device, a single sample can also be tested for apathogen in order to not only characterize the blood, but also pathogenscontained therein. In one embodiment, the moiety indicative of apathogen is selected from the group consisting of a pathogen, apathogen-associated antibody, a pathogen-associated nucleic acid, and apathogen-associated antigen.

In additional aspects, methods for testing a biological fluid (e.g.,blood) are provided. In the methods, the devices and systems providedherein are exposed to the fluid sample and the effect of the sample onthe device characteristics (e.g., resonant wavelength of a ringresonator) in response to binding events on the sample surface of thedevice are measured. Multiple devices can be used to compare the resultsof multiple different binding surfaces spread across multiple devices.The testing can be simultaneous across a plurality of devices in orderto facilitate real-time testing. Microfluidics can be used to deliverthe sample to the devices, as well as buffer and wash treatments.Automation can be accomplished by coordinating (e.g, by computer) thedelivery of sample to the device(s) and the measurement of the output ofthe device(s).

Exemplary Embodiment 1. Photonic Sensing in Undiluted Human Plasma

In the present study, we incorporated a novel self-adsorbingzwitterionic polymer, which is engineered to readily modify the nativeoxide of silicon-based devices for enhanced performance in complexbiological fluids. These zwitterionic polymers, composed ofcarboxybetaine methacrylate (CBMA) monomers, are charge dense, yetnet-neutral. This material property electrostatically induces surfacehydration as opposed to hydration by hydrogen bonding interactions,resulting in ultra-low protein fouling when exposed to human plasma andserum.

As demonstrated by the authors, the chemistry described herein haspreviously enabled the detection of biomolecules at clinically relevantsensitivities in complex media using plasmonic and mechanical sensors.Translating these zwitterionic chemistries to the microring resonatoropens the possibility of performing sensitive clinical assays on siliconphotonics. Using this approach, we report the label-free detection of aprotein at 10 ng ml⁻¹ in undiluted human serum, a first in the field ofsilicon photonics-based biosensing.

For this study, we utilized a microring resonator chip consisting of anumber of individual microring sensors. The microrings are rapidlyinterrogated by an external laser with a center frequency of 1560 nm(approximately 250 ms per microring). Real-time peak-fitting softwaredetermines the shift in resonance wavelength of the optical cavity as afunction of time. The biosensor chip is coated in a fluoropolymercladding to minimize waveguide losses. Portions of the cladding havebeen removed to expose the silicon oxide surface of the microringresonators for chemical modification and subsequent biosensingexperimentation. The remaining fluoropolymer-clad microring resonatorsserve as temperature and vibration reference controls (FIG. 1C). Mylarmicrofluidic gaskets are used to orthogonally address sets of microringresonators to perform chemical modifications and interrogate biologicalinteractions at the sensor surface. The geometry of the waveguide is a500×200 nm ridge, as shown in FIG. 1B, while the ring radius is 15 μm.

We introduced a self-adsorbing zwitterionic polymer DOPA-pCBMA (DpC)that dramatically reduces non-specific binding to the surface of thesilicon microring. Initially, we evaluated the ability of the DpCstrategy to prevent non-specific adsorption of protein to the oxidesurface on the silicon microring resonator. Microring resonatorbiosensors were exposed to solutions of DpC or bovine serum albumin(BSA) for surface passivation. The DpC coating was characterized byX-ray photoelectron spectroscopy (XPS) and ellipsometry. To assess theability of the biosensor surface coatings to resist non-specific proteinadsorption, we exposed DpC-modified resonators to solutions offibrinogen (FIG. 5) and undiluted human blood plasma (FIG. 2). FIG. 2illustrates relative resonance shift for three microrings with varyingsurface chemistries during exposure to undiluted human plasma. Asillustrated, microrings were subjected to buffer, then human plasma, andreturned to buffer. The net resonance shift at the 30 minute mark andbeyond is proportional to the amount of adsorbed material on the surfaceof the microring.

The net shift in resonance wavelength after exposure to human plasma wasused to assess the amount of ‘fouling’ protein adsorbed to the microringresonator surface. While BSA passivation of the sensor surface decreasedthe amount of ‘fouling’ protein by around 50% compared to the 660 μmshift seen on bare silicon oxide, the DpC coating resulted in only 5 μmof shift due to fouling in undiluted human plasma. This resultdemonstrates the noteworthy capability of DpC coatings to yieldultra-low fouling surfaces on silicon photonic devices.

In addition to effectively eliminating sensor response to biologicalnoise, the DpC strategy enables facile conjugation of capture ligandsthat can impart biological function to individual silicon microrings.When exposed to a sample solution, these capture element ligands bindtarget molecules generating a shift in microring resonance proportionalto the target analyte concentration. We employed carbodiimidechemistries to immobilize a monoclonal antibody (antiSA) specific forthe model protein analyte, streptavidin (SA). An isotype antibody(IgGi), sharing the same structure as antiSA, but with no bindingspecificity for SA, was immobilized to adjacent microring resonators toserve as a negative control during subsequent analyses (FIG. 6). Todemonstrate specific protein detection, phosphate-buffered saline (PBS)spiked with 20 μg/ml SA was flowed over the microrings, shown in FIG.3A. In FIG. 3A, the waveguide 105 and antiSA binding moiety 110 are usedto capture SA 120. As expected, the antiSA-DpC microrings exhibitedspecific binding of SA, while control IgGi-DpC microrings had nosignificant sensor response.

Further experiments were performed to explore the relationship of SAconcentration in PBS to the microring resonance shift. AntiSA-DpC andIgGi-DpC microrings were exposed to successively increasing SA-spikedPBS solutions, with intervening unspiked PBS washes, detailed in FIG.3B. The binding process of SA to antiSA-DpC is known to follow Langmuirstatistics and the relative resonance shift difference as a function oftarget analyte concentration can then be expressed as:

dλ=AαP/(1+αP)  (1)

Here, P is the concentration of SA in solution, α is the bindingcoefficient, A is a constant of proportionality depending on the numberof antiSA sites, and the amount of resonance shift per SA molecule, andfinally dλ is the relative resonance shift difference. A least-squaresfit of (1) to the peak shift differences between the antiSA-DpC andIgGi-DpC microrings seen in this experiment is shown in FIG. 3C.Langmuir statistics are followed, with a best-fit binding coefficient αof 0.0034 (ng ml⁻¹)⁻¹. The primary goal of this study was to demonstratethe detection of analyte in undiluted serum. Undiluted human serum wasspiked with SA at concentrations ranging from 10 ng ml⁻¹ to 10 μg ml⁻¹,encompassing a range of concentrations relevant to clinical diagnostics.Sensors were exposed to increasing concentrations of SA-spiked humanblood serum, with PBS buffer washes between samples, as seen in FIGS. 4A(detail) and 4B (full scan).

The results of several functionalized microring biosensors are shown.Large resonance shifts are seen on all rings, on the order of 650 μm,due to the difference in average refractive index between PBS and humanserum. A clear relative shift difference is seen between the antiSA-DpCmicrorings and the IgGi-DpC control microring, indicating that SA isbeing successfully detected. Some increase in the level of fouling isobserved when compared to the unfunctionalized DpC coated microringsshown in FIG. 2. This increase is due, in large part, to nonspecificbinding of serum proteins to the immobilized antibody capture ligands.Plotting the relative peak-shift differences for the antiSA-DpC andIgGi-DpC microrings, one recovers a relationship that follows Langmuirstatistics (FIGS. 7A and 7B). A least squares regression assumingrelation (1) yields a binding coefficient of 0.027 (ng ml⁻¹)⁻¹ and 0.026(ng ml⁻¹)⁻¹ for microrings 1 and 2, respectively, with the estimatednumber of binding sites 5.1×10⁵ and 3.8×10⁵. The difference in bindingsites between microrings is to be expected, as it is not possible toprecisely control the number of antibodies immobilized to the DpCscaffold on each device. Based on the approximate dimensions of the SAmolecule (5×5×5 nm³), the observed response at saturation corresponds to15% surface coverage of the microring in the case of antiSA-DpCMicroring 1. It is worth noting that we observe a difference in thebinding coefficients when characterizing antiSA-DpC/SA association inbuffer versus undiluted serum. This difference in binding coefficientswas anticipated due to the differences in solution composition (serumvs. buffer), and the results are comparable to what has been observed insimilar biosensing experiments.

The intrinsic noise of the measurements can be estimated by the squarederror between the fitted curve and the data, and is 2.8 μm and 2.5 μmRMS. Assuming a noise floor of 3σ, the functional biosensors exhibitsensitivities of approximately 10 ng ml⁻¹ of SA in undiluted humanserum. The achieved limit of detection is an order of magnitude moresensitive than the basic commercial colorimetric ELISA used for proteindetection in clinical assays. A further control experiment utilized asecond polyclonal antiSA antibody to confirm the specific capture of SAby the antiSA-DpC microrings and the absence of SA on the IgGi-DpCmicroring (FIG. 8).

To leverage the high sensitivity of silicon photonic-based sensingplatforms, robust chemical surface modification is imperative forlabel-free device performance in complex biological samples. For thefirst time, we have shown that label-free silicon photonic biosensorscan provide ELISA-like sensitivity with extraordinary selectivity inundiluted human serum. This study highlights a rapid, simple, andversatile chemical surface modification for silicon photonic biosensors:coatings of DpC can be deposited on sensors in minutes and can be usedto immobilize surface capture elements for biophotonic applications.This development in biocompatible silicon photonics represents asignificant step for the application of these devices in clinicaldiagnostics and the biomedical sciences.

Experimental Methods

Microring Resonator Biosensing Platform:

Silicon microring resonator biosensors and corresponding analysisinstrumentation were manufactured by Genalyte, Inc. (San Diego, Calif.).Each biosensor chip (6×6 mm) consists of an array of 32 individuallyaddressable microring resonators (30 μm in diameter) suitable forreal-time biosensing analysis. Twenty four of these microring resonatorsare exposed for biosensing and eight are coated with afluoropolymer-cladding to serve exclusively as temperature and vibrationreference controls. Exposed microring resonators can be surface modifiedusing chemistries that are compatible with the native oxide of thesilicon waveguides. To minimize waveguide losses, the entire chip iscoated with cladding with the exception of the exposed microringsensors, limiting all interactions with a biological sample todesignated resonators. An external cavity diode laser with a centerfrequency of 1560 nm rapidly interrogates grating couplers of individualmicroring resonators (approx. 250 ms per ring), measuring the shift inresonance wavelength of the optical cavity as a function of time.

Materials

All chemical reagents were purchased from Sigma-Aldrich, Corp (St.Louis, Mo.) and used without further purification unless otherwisenoted. Fraction V bovine serum albumin was purchased from EMD Chemicals(Gibbstown, N.J.). Fibrinogen (Fb; Fraction I, bovine plasma) waspurchased from Sigma-Aldrich. Human plasma and serum samples wereprovided by the Puget Sound Blood Center (Seattle, Wash.). Murinemonoclonal antibodies against streptavidin and monoclonal immunoglobulin(IgG1) control antibodies were purchased from Abcam, Inc (San Francisco,Calif.). Streptavidin and polycloncal anti-streptavidin antibodies (forsecondary antibody probe experiments) were purchased from VectorLaboratories (Burlingame, Calif.). All buffers used for biosensingexperiments were prepared using ultrapure deionized water (BarnsteadNanopure; Dubuque, Iowa). The pH of buffer solutions was adjusted using1M solutions of sodium hydroxide (NaOH) or hydrochloric acid (HCl).Phosphate-buffered saline (PBS, pH 7.4) was composed of 10 mM phosphate(1.9 mM KH₂PO₄, 8.1 mM Na₂HPO₄) and 150 mM sodium chloride (NaCl).Polymer deposition was performed in deposition buffer (10 mMtris(hydroxymethyl)aminomethane (Tris), pH 8.5). Surface grafted polymercoatings were activated using freshly prepared solutions of 0.4M1-Ethyl-3-[3-dimethylaminopropyl]carbodiimide hydrochloride (EDC) and0.1M N-hydroxysuccinimide (NHS) purchased from Sigma-Aldrich. Antibodyimmobilization buffer was composed of 10 mM HEPES(4-(2-hydroxyethyl)-1-piperazineethanesulfonic acid) at pH 7.8. Siliconwafers used for XPS characterization and ellipsometry measurements werepurchased from Silicon Valley Microelectronics (San Jose, Calif.).Hydrogen peroxide (H₂O₂) and sulfuric acid (H₂SO₄) solutions werepurchased from J.T. Baker (Phillipsburg, N.J.) and Sigma-Aldrich,respectively. Ethanol was purchased from Deacon Laboratories (King ofPrussia, Pa.).

Synthesis of DOPA-pCBMA (DpC) Conjugates

The synthesis of DpC polymer conjugates has been described in detail inprevious publications. Briefly, an initiator molecule was synthesizedcontaining two adhesive catechol groups (DOPA₂;N,N′-(2-hydroxypropane-1,3-diyl)bis(3-(3,4-bis(tert-butyldimethylsilyloxy)phenyl)-2-(2-bromo-2-methylpropanamido)propanamide))as described previously. Following synthesis of the CBMA monomer, DpCwas polymerized by atomic transfer radical polymerization (ATRP) methodsovernight. DpC conjugates were purified by dialysis, resulting in awhite powder. The tert-butyldimethylsilyloxy protecting groups weredeprotected using 1M tetrabutylammonium fluoride in tetrahydrofuran(THF). The polymer was washed extensively in THF, dried under reducedpressure, and was aliquoted for storage at −20° C. prior to use.

Polymer Deposition on Silicon Microring Resonator Arrays:

Prior to surface modification with DpC conjugates, microring resonatorbiosensor chips were vigorously cleaned to remove trace organics. Chipswere exposed to freshly prepared piranha solution (1:1 30% H₂O₂: 98%H₂SO₄) for 10 minutes with mild agitation to remove bubbles that formedon the fluoropolymer-clad chip surface (Caution! Piranha solution isextremely dangerous as it can react explosively in the presence oforganics). Biosensor chips were washed with copious amounts of waterprior to the deposition of DpC conjugates on the bare oxide surface ofsilicon microring resonators.

The deposition of DpC on microring resonator arrays was monitored inreal-time by analyzing the change in the resonance wavelength ofindividual microrings over time. All solutions were introduced tosensors at a flow rate of 20 μL/min using two alternating 50 μLnegative-pressure syringe pumps controlled by computer software(Genalyte, Inc). Prior to deposition, deprotected DpC conjugates werediluted to a concentration of 1 mg/ml in deposition buffer (10 mMTris-HCl; pH 8.5). The DpC solution was sonicated for 20 minutes toensure they were fully dissolved. The biosensor chip was exposed todeposition buffer to establish a signal baseline prior to introductionof the polymer solution. An array of microring resonators was exposed tothe sonicated DpC solution (1 mg/ml) for 15 minutes, resulting inhigh-density deposition of DpC conjugates on the exposed oxide of themicroring sensor surface. Microring resonators were then washed withdeposition buffer for 5 minutes, followed by an extensive wash (12minutes) in phosphate buffered saline (PBS; pH 7.4) to remove looselybound DpC conjugates. Finally, all sensors were exposed to depositionbuffer for the quantification of DpC surface coverage by comparing thenet shift in resonance wavelength to the initial sensor baselinedetermined in deposition buffer. As a control, a second array ofmicroring resonators was exposed to a solution of 1 mg/ml BSA prior toperforming protein fouling assays on each microring resonator biosensorchip. BSA surface functionalization was performed in parallel with DpCdeposition.

Protein Fouling Assays Using Protein Solutions:

The non-fouling character of DpC-coated and BSA-coated microringresonators was assessed by examining their resistance to protein foulingin simple and complex biological solutions. All solutions were flowed at20 μL/min and results were monitored in real-time as detailed above.After establishing a baseline in PBS, surface modified (DpC- andBSA-coated) microrings were exposed to 1 mg/ml fibrinogen for 30minutes, prior to returning to PBS for 30 minutes to assess dissociationof unbound protein (FIG. 5). The overall shift in resonance wavelengthwas used to compare the amount of protein fouling for DpC- andBSA-coated sensors. The ability of the surface modified sensors toresist protein fouling in complex biological solutions was assessedusing undiluted human plasma (FIG. 2). After establishing a baseline inPBS, sensors were exposed to undiluted human plasma for 15 minutes.After returning to PBS, the overall shift in resonance wavelength wasused to compare the amount of protein fouling for each surfacemodification strategy. The results were compared to the amount ofprotein fouling on unmodified (bare oxide) microring resonatorsfollowing extended exposure to undiluted human plasma.

X-Ray Photoelectron Spectroscopy (XPS) Characterization of DpC PolymerFilm:

Prior to XPS characterization of DpC films, the silicon substrate wascleaned with piranha solution, as detailed above. DpC was deposited onthe silicon from a 10 mM Tris-HCl buffer for 20 minutes. The substratewas rinsed vigorously with Tris-HCl buffer, PBS, ultrapure water, thendried under a stream of nitrogen for XPS analysis. XPS composition data(Table 1) were acquired on a Kratos AXIS Ultra DLD instrument equippedwith a monochromatic Al-Kα X-ray source (hv=1486.6 eV). XPS data werecollected at 0° takeoff angle in the hybrid mode with approximately 10nm sampling depth, using a pass energy of 80 eV. Three spots onduplicate samples were analyzed. Reported compositional data wereaveraged over multiple spots. Data analysis was performed on the CasaXPSsoftware (Casa Software Ltd.). As expected, DpC modification of thenative oxide of silicon decreases the percent composition of silicon (Si2p), while increasing the percent composition of organic material (C 1s,N 1s, O 1s) at the substrate surface.

TABLE 1 Relative compositions of bare Si and DpC-modified Si substratesas determined by XPS; n = 3, n/d: chemical species not detected. Bare Sichip DpC-modified Si chip Si 2p 45.9 ± 2.8% 11.1 ± 0.6%  C 1s 19.9 ±5.3% 56.1 ± 0.6%  O 1s 34.2 ± 2.5% 24.4 ± 0.8%  N 1s n/d 2.3 ± 0.1% Na1s n/d 2.8 ± 0.3% Cl 1s n/d 3.2 ± 0.3% Total 100% 100%

Ellipsometry Measurements of DpC Dry Film Thickness:

Prior to dry film thickness characterization, the silicon substrate wasexposed to solutions of DpC (1 mg/ml in 10 mM Tris-HCl) or BSA (1 mg/mlin PBS) for 30 minutes. Treated substrates were washed vigorously withtheir respective buffers, then ultrapure water, and dried under a streamof nitrogen. To determine the dry film thickness of DpC and BSA films onsilicon, measurements were obtained using an ellipsometer (M-2000; J.A.Woollam, Inc; Lincoln, Nebr.). The amplitude component (Ψ) and phasedifference (Δ) were measured for DpC-modified and BSA-coated siliconsubstrates from 200-1000 nm wavelengths at varying angle of incidence(65°, 70°, 75°) at 9 spots on each chip. The data obtained frommeasurements was fit using a generalized Cauchy layer (A_(n)=1.45,B_(n)=0.01, C_(n)=0). The thickness of the native oxide layer of thesilicon substrates was determined as described above and applied to theCauchy layer fit (˜2.2 nm). The average dry film thickness of DpC andBSA films on the silicon substrates (Table 2) suggest that the polymersurface modification strategy results in very thin coatings capable ofstrongly resisting protein fouling.

TABLE 2 Ellipsometry characterization of DpC and BSA films on nativeoxide of Si substrates Average Thickness (n = 9) MSE BSA film 3.27 ±0.25 nm 3.16 DpC film 1.08 ± 0.10 nm 2.83

Immobilization of Antibodies to DpC-Modified Microring ResonatorBiosensors:

The terminal carboxylate groups of DpC-coated microring resonators wereactivated using carbodiimide chemistry prior to antibody immobilization.A freshly prepared activation buffer (0.4M EDC, 0.1M NHS) was passedover the DpC-modified sensors twice for 5 minutes, separated by a 1minute wash steps with ultrapure water. Activated DpC-coatings onmicrorings were exposed to either monoclonal anti-streptavidin (antiSA)or immunoglobulin control antibodies (IgGi, negative control) asfollows. Activated DpC-microrings were briefly exposed to immobilizationbuffer (10 mM HEPES, pH 7.8) to establish a baseline, followed byimmediate exposure to antibody solutions (20 μg/ml in 10 mM HEPES, pH7.8) for 12 minutes (FIG. 6). After antibody immobilization, theactivated-DpC was quenched via hydrolysis at elevated pH (10 mM HEPES,300 nM NaCl, pH 8.2) for 15 minutes, followed by immobilization bufferto reestablish baseline and determine the amount of immobilizedantibody.

Selection Criteria for Functional Microrings for Further Analysis:

Small differences in the level of immobilized antibody, as determined bythe overall shift in resonance wavelength, had a significant affect onthe performance of functionalized sensors. Therefore, we analyzedmicroring resonators that responded with 200 μm of overall resonanceshift following antibody immobilization for analyte detection inundiluted human serum. We also found that high levels of antibodyimmobilization (>200 μm in our studies) led to a decrease in thenon-fouling nature of the DpC-coating in complex media, resulting inincreased non-specific adsorption and a decreased capacity to detectanalyte at low concentrations. We found that this quantitative cut-offwas not required for detection of analyte in buffer for thisproof-of-concept demonstration. This process emphasizes the need tofabricate and analyze arrays of silicon photonic devices in order toassess the uniformity of elaborate surface modifications.

Detection of Streptavidin in Buffer and Undiluted Human Serum:

The model protein streptavidin (SA) was used to demonstrate specificdetection of analyte in either buffer or undiluted human serum byantibody functionalized DpC-modified microring resonators. IgGi-DpCmodified microrings served as a negative control for non-specific SAbinding to immobilized antibodies. To demonstrate analyte detection inbuffer, SA was diluted in PBS at concentrations ranging from 50 ng/ml to10 μg/ml. After establishing a signal baseline in PBS, increasingconcentrations of SA in PBS were introduced to microring resonators for5 minutes. Concentration steps were separated by 5 minute buffer (PBS)washes. The shift in resonance wavelength of antiSA-DpC microrings wascompared to IgGi-DpC microrings to demonstrate specific streptavidindetection in buffer.

SA detection in undiluted serum was confirmed by exposing microringresonators to a series of SA-spiked human serum samples. To obtainsensitive device measurements, undiluted SA-spiked human serum samples(10-10,000 ng/ml) were flowed over microrings for 10 minutes per sample(20 μL/min) using a negative-pressure syringe pump (Chemyx, Inc;Stafford, Tex.) equipped with 5 mL glass syringes (Hamilton, Co; Reno,Nev.). Samples were separated by 5 minute buffer (PBS) washes. SpecificSA binding was defined as the difference in sensor response betweenantiSA-DpC and IgGi-DpC microrings. Functionalization of surfacegrafted-DpC resulted in a small increase in protein fouling compared tounmodified DpC coatings. However, there is little protein accumulationover time. Specifically, antibody functionalized DpC-modified microringsshowed ˜50 μm of fouling after an initial exposure to serum. However,over 60 minutes of additional exposure to undiluted serum resulted in˜55 μm of additional protein fouling. These results demonstrate that DpCcoatings may be functionalized with biomolecules while largely retainingtheir non-fouling properties during extended exposure to undiluted humanserum.

Resonance Shift Calculations:

Consider a cross-section of a waveguide, expressed as c(x,y) over areaΩ. It can be shown that if a portion of the waveguide cross section, Ω′,experiences a shift in dielectric constant dc(x,y), the shift ineffective index (for small changes) will be:

$\begin{matrix}{{dn}_{eff} = \frac{\left. {\int\int_{\Omega^{\prime}}} \middle| E \middle| {}_{2}{{{dɛ}\left( {x,y} \right)}{dxdy}} \right.}{2Z_{O}{\int{\int_{\Omega}{{{Re}\left( {E^{*} \times H} \right)}{\square{zdxdy}}}}}}} & (2)\end{matrix}$

For the experiments described in this study, two types of index shiftsare relevant. First, an index shift can be created by a bulk change inthe refractive index surrounding the waveguide, perhaps from switchingthe fluid from PBS to human serum. For the 500×200 nm waveguideutilized, if the cladding index is close to 1.35, the index ofphosphate-buffered saline (PBS), this shift can be calculated as:

dn _(eff)=0.06dε _(clad)  (3)

Another type of index shift is caused by a molecule becoming bound tothe surface of the microring. In this case, the integral in (2) shouldbe taken over the entire area into which the molecule might bind, callthis Ω′, and then the result discounted by A/Ω′, where A is thecross-sectional area of the molecule. The molecules under investigationin this work have typical sizes ranging from 5 to 15 nm. The hydratedDpC layer is also estimated to be on the order of 10 nm. Therefore allindex shifts occur within 30 nm of the surface of the waveguide. Theoptical fields exhibit minimal falloff over such a short distance,rendering the index shift independent to small changes in positioning ofthe molecules. Once the index shift is known, it is easy to calculatethe shift in resonance wavelength:

$\begin{matrix}{{d\lambda} = {\lambda \frac{{dn}_{eff}}{n_{g}}}} & (4)\end{matrix}$

Here n_(g) is the group index of the waveguide, 4.05 near 1550 nm. Thewaveguide effective index is 2.33. For a molecule, (4) would need to bemultiplied by L′/L, the length of the molecule, divided by thecircumference of the microring, in this case 94 μm. Combining all theseexpressions, we have:

$\begin{matrix}{{d\lambda} = {\left( \frac{\lambda}{n_{g}L} \right)\left( \frac{\frac{1}{\left. {\Omega^{\prime}{\int\int_{\Omega}^{\;^{\prime}}}} \middle| E \middle| {}_{2}{dxdy} \right.}}{2Z_{O}{\int{\int_{\Omega}{{{Re}\left( {E^{*} \times H} \right)}{\square{zdxdy}}}}}} \right)\frac{m}{\rho}\left( {ɛ_{molecule} - ɛ_{background}} \right)}} & (5)\end{matrix}$

Here m is the total mass of bound molecules, ρ is the density of themolecule, and ε_(molecule) and ε_(background) are the dielectricconstants of the molecule and background materials, respectively. Wehave due to numerical calculation:

$\begin{matrix}{\left( \frac{\frac{1}{\left. {\Omega^{\prime}{\int\int_{\Omega}^{\;^{\prime}}}} \middle| E \middle| {}_{2}{dxdy} \right.}}{2Z_{O}{\int{\int_{\Omega}{{{Re}\left( {E^{*} \times H} \right)}{\square{zdxdy}}}}}} \right) = {0.93\frac{1}{{\mu m}^{2}}}} & (6)\end{matrix}$

Also:

$\begin{matrix}{\left( \frac{\lambda}{n_{g}L} \right) = 0.0041} & (7)\end{matrix}$

We can now calculate the index shift per binding event, or equivalently,the total mass bound, for a molecule where we know the refractive indexand density. As noted in the main paper, there is a simple linearrelation between the bound mass and the wavelength shift of a resonancepeak. Streptavidin (SA) and antiSA should have approximate refractiveindex 1.45, density 1.35 g/ml, and respective molecular weights of 60,150 kDa. We note that the approximately 600 μm resonance shift seen inthe human serum experiments suggests that the index of human serum is1.36. The shifts predicted in (4) should then be nearly identical for agiven bound molecule, regardless of whether the background is PBS orhuman serum. We have, finally, 7.86×10⁻⁵ μm of shift per SA bindingevent in PBS. This can also be expressed as a mass sensitivity; 0.79μm/fg of resonance shift is expected, in agreement with the 0.89 μm/fgmeasured elsewhere for a similar system. The number of SA binding sitesfor a given peak resonance shift difference can then be readilycalculated.

Confirmation of Specific Streptavidin Binding to antiSA-DpC Microrings:

To confirm specific SA binding to antiSA-DpC microrings, a polyclonalantiSA antibody (50 μg/ml) was flowed over sensors after detection of SAin serum. By analyzing the formation of an antibody “sandwich”(antiSA/SA/antiSA-DpC), we were able to determine if there was specificcapture of SA by the antiSA-DpC sensors. The antiSA-DpC microringsexhibited a significant peak shift, while the IgGi-DpC control microringshowed no shift (FIGS. 7A and 7B). This suggests that the bound SA onthe antiSA-DpC microring was able to bind the polyclonal antiSAantibodies as expected, while simultaneously confirming a lack of SAbound to the IgGi-DpC microring. The net peak shift, approximately 100μm, was around a factor of 3 larger than the final relative peak shiftof the antiSA and IgGi microrings, which was approximately 30 μm. Thisis as expected, due to the relative molecular weights of 60 and 150 kDafor SA and antiSA respectively. Further, this result indicates anapproximate 1:1 correspondence between the final number of SA moleculesbound, and the final number of polyclonal antiSA antibodies bound fromsolution.

Exemplary Embodiment 2. Hierarchical Anti-Fouling Layer for PhotonicSensing

Surface chemistries for biosensors, implantable medical devices targeteddrug/gene delivery carriers, tissue scaffolds, and targeted molecularimaging probes in complex media remain a great challenge due to highnonspecific adsorption and low binding capacity of molecular recognitionelements. Currently, few materials have been developed to reducenonspecific protein adsorption, including poly(ethylene glycol) (PEG),mannitol tetraglyme, and zwitterionic polymers. The effectiveness ofprotein resistant materials relies on their high surface packingdensities. Unfortunately, highly dense two-dimensional (2D) polymerfilms elicit the limitation of a low ligand-binding capacity. At thesame time, a three-dimensional (3D) carboxymethylated dextran-basedhydrogel binding matrix was previously developed, enabling very highprotein loading due to an open polymer structure. However, this openstructure only provides weak surface resistance to nonspecific proteinadsorption, particularly in complex media such as blood. New polymerswith precisely controlled architecture are desirable for exploring novelstructure-property correlations and achieving unique properties for manyapplications. In this communication, we propose a unique strategy ofdeveloping hierarchical polymer films with structurally regulatedfunctionalities through integrating 2D and 3D structures so as toachieve ultra low nonspecific binding and high loading of molecularrecognition elements.

Our first attempt to construct a binding platform with a hierarchicalarchitecture was demonstrated with two distinct surface-initiatedtechniques. These “grafting from” approaches, based on controlled“living” radical polymerizations, were particularly promising for thepreparation of polymer brushes as they permit precise control overchemical composition, film thickness, and architecture. As shown inScheme 1, the films were prepared via surface initiated atom transferradical polymerization (SI-ATRP) and surface initiatedphotoiniferter-mediated polymerization (SI-PIMP). The first layer wasgrown in a controlled manner to reach a high surface packing density.The second layer, with a low surface packing density, was achievedthrough “termination” or “regeneration” of the living capped species atthe polymer chain end for SI-ATRP and SI-PIMP, respectively.

Due to the dual functionality of poly(carboxybetaine) (pCB) films, aproof-of-concept experiment was performed with zwitterionic pCB using asurface plasmon resonance (SPR) biosensor for demonstrating the novelhierarchical architecture. Previous reports of surface-tethered pCBbrushes formed by both SI-ATRP and SI-PIMP have achieved excellentresistance to nonspecific protein adsorption in the presence of complexmedia, such undiluted human blood serum and plasma, to fouling levelsbelow 5 ng cm⁻². These fouling levels can be maintained following theimmobilization of around 250 ng cm⁻² of antibody using conventional1-ethyl-3-(3-dimethylaminopropyl)-carbodiimide and N-hydroxysuccinimide(EDC/NHS) coupling chemistry under biologically friendly conditions.However, this functionalization level only corresponds to an IgGmonolayer. Herein, pCB-based platforms with hierarchical structures on aSPR sensor surface are presented for the sensitive quantification of IgGimmobilization, antigen binding, and nonspecific protein adsorption.

As shown in FIG. 9, ATRP was combined with a “termination” approach fordemonstrating the novel architecture. ATRP involves a dynamicequilibrium between activated propagating radicals and dormant halideend-capped polymer chains, yielding low polydispersity and controlledgrowth. This “living” characteristic enables re-initiation frommacroinitiators for the synthesis of block copolymers. In order toachieve hierarchical pCB films possessing ultra low fouling and highloading properties via SI-ATRP, a densely packed first layer was grownfrom a gold coated SRP chip modified with an alkyl bromide terminatedself-assembled monolayer (SAM). The chips were then submerged in amethanolic solution containing 2,2′ bi-pyridine, CuBr, and CB monomerunder nitrogen protection and allowed to react overnight. The resultingthickness was 7.6±0.3 nm (Table 3). Importantly, these conditionsenabled a highly dense yet thin film to be grown. While the high densityis key for achieving low fouling, a thin film is desired for manysensing applications, such as SPR, as the signal intensity/sensitivitydecays exponentially from surface of the metal substrate. To establish ahierarchically structured pCB film for increasing the binding capacity,the macroinitiator density for re-growth of the second pCB layer wasregulated via azide substitution of bromide species thus “terminating”the future growth of the corresponding chains during the second ATRPreaction. The density of the polymer chains can be controlled by theazide concentration and reaction time. In this study, a 2 hr submersionusing an azide concentration of 0.1 M produced the optimal second layerpolymer density for protein immobilization. For the growth of the secondpCB layer, water-accelerated polymerization with a solvent consisting of50% water in methanol was employed to induce a high polydispersity ofpolymer chains. The resulting thickness of the structured film withazide substitution was higher than that without treatment. This is inagreement with previous reports showing rapid bimolecular termination athigh initiator densities using aqueous ATRP whereas more diluteinitiators enabled continued linear and controlled polymer growth.

TABLE 3 Thickness of films prepared via SI-ATRP and SI-PIMP with andwithout treatment to capped species. SI-ATRP SI-PIMP Thickness Thickness(nm) (nm) First Layer 7.6 ± 0.3 10.8 ± 0.8 Re-growth (without treatment)13.2 ± 0.3  32.1 ± 0.6 Re-growth (with treatment) 17.5 ± 0.9  46.1 ± 1.6

In contrast to SI-ATRP, SI-PIMP is apt to release capped species fromthe polymer chains during polymerization, primarily due to bimoleculartermination. This irreversible termination significantly hampers futurepolymer growth. Therefore, in order to control the chain density of thesecond layer via SI-PIMP, a “regeneration” approach was adopted in whichthe addition of a deactivator, tetraethylthiuram disulfide (TED), wasable to preserve the end-capped photoiniferter groups on the graftedpolymers for re-growth of the second layer with controlled graftingdensity. SPR gold substrates were first modified with the photoiniferter(N,N-(Diethylamino)-dithiocarbamoylbenzyl(trimethyoxy)-thiol (DTCA)) toform SAMs. Similarly to SI-ATRP, the first layer for SI-PIMP was alsosynthesized in 100% methanol to form a highly dense and thin film.Reactions were conducted using a 30 min UV irradiation and the resultingfilm thicknesses are shown in Table 3. The first layer thicknessesprepared with 2 μM TED was comparable to that without TED (11.1±0.6 nm).Subsequently, the films were re-initiated in a 90% water/methanolsolution resulting in TED treated films with greater thicknesses thanthose made without, reflecting the ability of TED for preserving thereactive photoiniferter end groups and thereby maintaining the “living”characteristic of SI-PIMP.

In this study, functionalization and fouling tests were monitored insitu using a custom-built SPR sensor with wavelength modulation. Forantibody immobilization, the films were activated using EDC/NHS couplingchemistry followed by injecting an anti-human thyroid stimulatinghormone (anti-TSH) IgG solution. The unreacted NHS esters were thenhydrolyzed back into the original carboxylate groups using 10 mM sodiumcarbonate buffer with 300 mM NaCl at pH 10. As shown in FIG. 10A, thefunctionalization levels were estimated as 195.9 ng cm⁻² and 417.0 ngcm⁻² for one-layer (“One” in FIGS. 10A-10C) and hierarchical (“HA” inFIGS. 10A-10C) films prepared via SI-ATRP; 253.0±14.8 and 792.7±54.7 ngcm⁻² for one-layer and hierarchical films from SI-PIMP. Here, anincrease in binding capacities for IgG molecules on pCB films wasobserved for the hierarchical architecture. For the IgG functionalizedone-layer films, the binding capacities were similar to that obtainedwith carboxyl-terminated SAMs. Although pCB provides abundant carboxylgroups for biomolecule conjugation, highly-packed polymer brushes hamperthe penetration of molecules due to steric hindrance and thereforemodification only takes place at the accessible functional groups on thetopmost layer of the pCB films. However, for the hierarchical films, thechain densities of the second layer were controlled via the terminationand regeneration approaches. Constructed from highly dense first layers,the loose second layers allowed diffusion of antibodies thus enablingconjugation with NHS esters throughout the entire second layer. Thecontrol experiments using the structured films without treatment ofcapped species were also conducted. IgG immobilization levels werereduced by 30% and 64% for SI-ATRP and SI-PIMP, respectively, comparedto the corresponding treated hierarchical pCB films. This evidenceindicates that a sufficient number of accessible binding groups forprotein modification, made apparent by the larger second-layer filmthicknesses of the treated films, are a determining factor of the ligandloading capacity.

FIG. 11 is an SPR sensorgram for the fouling test in the presence ofundiluted blood plasma or serum on hierarchical pCB films. FIG. 12 is anSPR sensorgram following EDC/NHS activation, the pCB films were modifiedwith TSH IgG, deactivated with SC buffer, and then used for TSH antigendetection.

The protein fouling levels on one-layer films and hierarchical pCB filmsbefore and after IgG functionalization were tested by flowing undilutedblood serum or plasma (FIG. 10B). All fouling levels were very low as aresult of the highly-packed first pCB layers serving as ultra lowfouling backgrounds. As a comparison, the fouling level for a loose pCBone-layer film with a thickness of 12.3 nm prepared from 50% water inmethanol by SI-PIMP was as high as 54.3 ng cm⁻² in the presence ofserum. These results indicate that the high performance of pCB foreffective resistance against non-specific adsorption and high ligandloading is established on the basis of control over the polymerarchitecture.

Solutions containing TSH antigen were flowed over the functionalizedsurfaces to evaluate the antigen detection ability (FIG. 10C) and thebio-activities (molar ratios of antigen to antibody) of the bindingplatforms. TSH binding capacities were 42.1±2.4 ng cm⁻² and 128.5±26.2ng cm⁻² for one-layer and hierarchical pCB films from SI-PIMP,respectively. The corresponding bio-activities were 0.89 and 0.87. Theone-layer and hierarchical pCB films made via SI-ATRP bound 30.4 ng cm⁻²and 74.7 ng cm⁻² of TSH, respectively, with bio-activities of 0.80 and0.93. Thus, this study demonstrates that the binding capacity forantigens is well correlated to the degree of antibody immobilization andthat the bio-activity of the film is not affected by the pCBhierarchical architecture.

In conclusion, new methodologies were developed for creatingsurface-initiated polymer brushes with hierarchical architecturespossessing distinctive structurally regulated functionalities. Throughcombining the benefits of 2D and 3D polymer structures, a unique bindingplatform with ultra low fouling and high loading was established toserve as a new model for advancing surface chemistry needs. While thiswork was demonstrated with pCB prepared via the “grafting from” approachusing a SPR biosensor, this is a powerful and yet generic concept whichcan be applied to other surface chemistries, such as dextran orPEG-based materials, and is easily adaptable to other sensing platformsand devices. Furthermore, the preparation method for the hierarchicalarchitecture can be modified for other applications, such as “click” orthiol-ene “grafting to” chemistries. This research illustrates the greatpromise of this hierarchical surface coating development formulti-functional surface chemistry in biotechnological andnano-engineering applications.

Experimental Methods

pCB Films Via SI-ATRP

Mercaptoundecyl bromoisobutyrate (SI-ATRP initiator) and carboxybetaineacrylamide (CB) monomer were synthesized as described previously. SAMson cleaned SPR chips of ATRP initiator were formed by soaking overnightin ethanol (0.1 mM). Upon removal, the chips were rinsed with ethanol,THF, ethanol, and then dried and placed in a custom glass tube reactorunder nitrogen. In a separate glass tube, CuBr (8.86 mg),2,2′-bipyridine (57.85 mg), and CB (600 mg) were added and placed undernitrogen. The solids were dissolved in nitrogen purged methanol (4 mL)and transferred to the chips and reacted for 24 hours at 25° C. in ashaker bath. For single layer films, the chips were rinsed with waterand stored overnight in PBS. For hierarchical films, the solution wasquenched with CuBr₂ (275.87 mg) in methanol (4 mL) and then rinsed withmethanol, water, and submerged in PBS. The second block was then grownvia repeating the above procedure but using a nitrogen purgedmethanol:water (1:1) and reacting for 3 hours. Termination of brominegroups and replacement with non-reactive azide moieties for reducing thesecond block polymer density was achieved by submerging the single layerchips in an aqueous azide solution (0.1 M) for 2 hours, removing andrinsing with PBS, water, and then drying for ATRP.

pCB Films Via SI-PIMP

The DTCA photoiniferter was synthesized as described previously. SAMs oncleaned SPR chips of the photoiniferter were formed by soaking overnightin THF containing DTCA (2 mM) followed by rinsing with THF and dryingwith a stream of air. For single layers, the photoiniferter modifiedchip was transferred to a quartz reaction tube along with 170 mg of CBmonomer and placed under nitrogen. Nitrogen purged methanol (5 mL)containing TED (2 μM) was transferred to the reaction tube. Thephoto-polymerization was then conducted for 30 min using a UV lamp (302nm) coupled with a 280 nm cutoff filter for preventing deterioration ofthiol-gold bonds. Following the reaction, the chips were removed andrinsed with water, PBS, and then submerged in PBS. For the hierarchicalfilms, the single layer film was re-initiated using the identicalprocedure except for the using nitrogen purged methanol:water (10:90) inthe absence of TED.

Ellipsometry

The thickness of the pCB films were determined using an ellipsometer(Model alpha-SE, J.A. Woollam, Lincoln, Nebr.) using the 380-900 nmwavelength range at an Incidence Angle of 70°. The Results were Fittedto a Cauchy Module.

Non-specific Protein Adsorption, Antibody Modification, and AntigenDetection

The non-specific adsorption, antibody immobilization, and antigendetection was monitored using a custom-built four-channel SPR sensorwith the Kretschmann configuration and wavelength modulation asdescribed previously. SPR chips were made of a glass slide coated withtitanium (2 nm) followed by gold (48 nm) using an electron beamevaporator. A 1 nm SPR wavelength shift corresponded to a change in theprotein surface coverage of 17 ng cm⁻² which was corrected to accountfor loss of sensitivity due to the polymer films using previouslydescribed methods. For fouling experiments, undiluted human serum orplasma were injected (10 min, 40 μL min⁻¹) and the wavelength shiftbetween PBS baselines was converted to a surface coverage. Anti-TSH wasimmobilized by first injecting 10 mM sodium acetate (SA, pH 5) followedby EDC/NHS (0.2 M/0.05 M in water) for 7 min at 30 μL min⁻¹. Anti-TSH(50 μg mL⁻¹ in 10 mM HEPES pH 7.5) was injected (20 min, 20 μL min⁻¹)followed by deactivating with 10 mM sodium carbonate containing 300 mMsodium chloride (pH 10) for 10 min and the SA both at 30 μL min⁻¹.Immobilization was calculated as the difference between SA baselinesbefore IgG injection and after deactivation. TSH was antigen binding wasthen monitored by first injecting PBS and then antigen (1 μg mL⁻¹ in PBSat 40 μL min⁻¹) following by PBS.

Materials

Copper (I) Bromide (99.999%), 2,2′-bipyridine (BPY, 99%),tetrahydrofuran (THF), Tetraethylthiuram disulfide (TED), methanol,4-(2-hydroxyethyl)-1-piperazineethanesulfonic acid (HEPES), andphosphate buffered saline (PBS, 0.01 M phosphate, 0.138 M sodiumchloride, 0.0027 M potassium chloride, pH 7.4) were purchased fromSigma-Aldrich (St. Louis, Mo.). Ethanol (200 Proof) was purchased fromDecon Laboratories (King of Prussia, Pa.). Sodium carbonate anhydrouswas purchased from EMD Chemicals (Darmstadt, Germany). Sodium chloride(NaCl) and ether were purchased from J.T. Baker (Phillipsburg, N.J.).Sodium acetate anhydrous was purchased from Fluka (subsidiary of SigmaAldrich, St. Louis, Mo.). 1-Ethyl-3-(3-dimethylaminopropyl) carbodiimidehydrochloride (EDC) and N-hydroxysuccinimide (NHS) were purchased fromAcros Organics (Geel, Belgium). Pooled human serum was purchased fromBiochemed Services (Winchester, Va.). Antibody to thyroid stimulatinghormone (anti-TSH) and the TSH antigen were purchased from ThermoFisherScientific (Waltham, Mass.). Water used in the experiments was purifiedusing a Millipore water purification system with a minimum resistivityof 18.2 MΩ cm.

Exemplary Embodiment 3. Blood Typing

Direct Typing

To prepare biosensor chips for direct red blood cell (RBC) typingassays, microring resonator arrays were coated with adsorbed Protein A.After briefly washing the chip, monoclonal anti-blood group IgGantibodies (anti-A and anti-B) as well as a negative control monoclonalIgG (antiPSA) were microspotted on separate regions of each chip. Theentire biosensor chip was blocked with bovine serum albumin to minimizenon-specific adsorption of red blood cell fragments. The membranes ofred blood cells (Type A or Type B RBCs) were ruptured by resuspendingcells in a hypotonic lysis buffer, and cell fragments were washed viacentrifugation to remove expelled cellular contents including hemoglobinfrom solution. Ruptured red blood cells were diluted inphosphate-buffered saline and introduced to functionalized sensor arraysat a flow rate of 20 μl/min for 15 minutes. The sensor responses ofanti-A and anti-B functionalized microring resonators were normalized tothe response of the negative control antiPSA sensors. The results of aplurality of anti-A and anti-B sensors are presented in FIG. 13 (testingfor Type A RBC) and FIG. 14 (testing for Type B RBC). The specificdetection of both Type A and Type B RBC is apparent from the discloseddata.

Indirect Typing

Arrays of microring resonators were coated with the non-fouling DpCzwitterionic polymer as previously described herein. Surface grafted-DpCpolymer chains were chemically modified to covalently immobilize thestreptavidin (SA) protein for further derivatization of the sensors withbiotinylated capture elements. SA-DpC microrings were microspotted withbiotinylated blood group antigens (A and B antigens) for specificcapture of anti-blood group antibodies from human sera. Functionalizedsensor arrays were then exposed to undiluted human plasma (Type A orType B plasma), and the differential sensor response was compared todetermine the presence of anti-blood group antibodies in blood samples.The testing of sensors exposed to type A plasma is illustrated in FIG.15. A summary of test data from multiple sensors exposed to both type Aand type B plasma is illustrated in FIG. 16. The consistency of theresults presented in FIGS. 15 and 16 demonstrate the usefulness of thephotonic devices to determine blood type.

While illustrative embodiments have been illustrated and described, itwill be appreciated that various changes can be made therein withoutdeparting from the spirit and scope of the invention.

The embodiments of the invention in which an exclusive property orprivilege is claimed are defined as follows:
 1. A photonic device fordetermining blood type or immune sensitization to blood type,comprising: a sample waveguide having a sample surface; and a bindingcoating covering and being in optical communication with at least aportion of the sample surface, the binding coating being configured tobind a target moiety indicative of blood type or immune sensitization toblood type, wherein the photonic device is configured such that lightpassed through the sample waveguide has an evanescent field that extendsa distance beyond the sample waveguide sufficient to detect the targetmoiety indicative of blood type.
 2. The photonic device of claim 1,wherein the target moiety indicative of blood type is attached to ablood cell.
 3. The photonic device of claim 2, wherein the blood cell isselected from the group consisting of a white blood cell, a red bloodcell, a platelet, and a microparticle.
 4. The photonic device of claim1, wherein the binding coating is an antigen and wherein the targetmoiety indicative of blood type is an antibody indicative of blood orpathogen antigens.
 5. The photonic device of claim 1, wherein thebinding coating is antifouling.
 6. The photonic device of claim 1,wherein the binding coating is zwitterionic.
 7. The photonic device ofclaim 6, wherein the binding coating comprises a poly(carboxybetaine).8. The photonic device of claim 1, wherein the binding coating comprisesa plurality of layers, including an antifouling layer and a capturelayer.
 9. The photonic device of claim 8, wherein the antifouling layeris zwitterionic.
 10. The photonic device of claim 8, wherein the capturelayer comprises at least one binding moiety.
 11. The photonic device ofclaim 10, wherein the binding moiety is configured to bind to the targetmoiety indicative of blood type.
 12. The photonic device of claim 8,wherein the antifouling layer is bound to the capture layer.
 13. Thephotonic device of claim 1, wherein the binding coating is covalentlyattached to the sample waveguide.
 14. The photonic device of claim 1,wherein the binding coating is not bound to the sample waveguide. 15.The photonic device of claim 1, wherein the binding coating has a firstrefractive index prior to binding the target moiety indicative of bloodtype and a second refractive index after binding the target moietyindicative of blood type, wherein the first refractive index and thesecond refractive index are different, and wherein an evanescent fieldof electromagnetic radiation of a first wavelength extends beyond thebinding coating and into any binding moiety bound to the bindingcoating.
 16. The photonic device of claim 1, wherein the samplewaveguide is a portion of a photonic device selected from the groupconsisting of a resonator and an interferometer.
 17. The photonic deviceof claim 1, wherein the sample waveguide is a portion of a photonicdevice selected from the group consisting of a ring resonator, a Braggreflector, and a Mach-Zehnder interferometer.
 18. A photonic system fordetermining a blood type, comprising: (1) a first photonic device fordetermining the blood type, comprising: a first sample waveguide havinga first sample surface; and a first binding coating covering and beingin optical communication with at least a portion of the first samplesurface, the first binding coating being configured to directly bind toa blood cell target moiety indicative of the blood type, wherein theblood cell target moiety is attached to a blood cell body selected fromthe group consisting of a blood cell, a blood cell membrane, a bloodcell fragment, a microvesicle, and a blood cell-associated antigen, andwherein the first photonic device is configured such that light passedthrough the first sample waveguide has an evanescent field that extendsa distance beyond the first sample waveguide sufficient to detect thebound blood cell target moiety; and (2) a second photonic device fordetermining blood type, comprising: a second sample waveguide having asecond sample surface; and a second binding coating covering and beingin optical communication with at least a portion of the second samplesurface, the second binding coating being configured to bind to anantibody indicative of the blood type, wherein the second photonicdevice is configured such that light passed through the second samplewaveguide has an evanescent field that extends a distance beyond thesecond sample waveguide sufficient to detect the antibody indicative ofthe blood type.
 19. The photonic system of claim 18, wherein the firstphotonic device and the second photonic device are configured tosimultaneously determine the blood type.
 20. The photonic system ofclaim 18, further comprising a computer configured to determine theblood type using output from both the first photonic device and thesecond photonic device.
 21. The photonic system of claim 18, furthercomprising a reference waveguide that does not have any binding coating.22. The photonic system of claim 18, further comprising a third photonicdevice configured to bind a target moiety indicative of a pathogen. 23.The photonic system of claim 22, wherein the target moiety indicative ofa pathogen is selected from the group consisting of a pathogen, apathogen-associated antibody, a pathogen-associated nucleic acid, and apathogen-associated antigen.